Poly (diol co-citrate) hydroxyapatite composite for tissue engineering and orthopaedic fixation devices

ABSTRACT

The present invention is directed to a novel poly (diol citrates)-based bioceramic composite materials created using completely biodegradable and a bioceramic material polymers that may be used in implantable devices. More specifically, the specification describes methods and compositions for making and using bioceramic composites comprised of citric acid copolymers and a bioceramic material.

The present application claims benefit of U.S. Provisional ApplicationNo. 60/771,241 filed Feb. 8, 2006. The entire text of the aforementionedapplication is incorporated herein by reference.

FIELD OF THE INVENTION

The present invention describes new composites for use in orthopedicdevices.

BACKGROUND

Orthopaedic, cranio-facial, and oral-maxillofacial surgeons often usetissue fixation devices such as pins, plates, and screws that are madefrom poly-1-lactide (PLLA),a biodegradable polymer[1-7]. Althoughbiodegradable devices can have significant advantages over their metalcounterparts, there are concerns with their use. These include slowdegradation, which can be as long as 5 years, and their inability tofully integrate with bone, which can be a problem for revision surgeries[8,9]. Also, PLLA bone screws can fracture during the fixationprocedure. A strategy to improve the osteointegration capacity of PLLAhas been to blend it with hydroxyapatite (HA), a bioceramic that can befound in natural bone mineral. Although HA is very brittle and hard toprocess into fixation devices of sufficient strength and fatigueresistance, it can impart osteoconductivity to polymers [10, 11].Researchers have shown that under certain conditions, addition of HAparticles can improve the mechanical properties of the polymer componentwhen used in a composite blend [12-14]. Therefore composites of polymerswith bioceramics may be a suitable compromise to meet mechanicalproperty requirements and achieve osteointegration of the implant.Nevertheless, there remains a significant problem in that the PLLAcontinues to slowly degrade over a period of time. Moreover,incorporation of more than about 30 wt. % of HA into any such compositeleads to a material that is too brittle for use in implantable devices.Thus, there remains a need for composite materials that arebiocompatible, can be easily processed and will fully integrate with thesurrounding bone and tissue within a year of implantation.

SUMMARY OF THE INVENTION

The present invention provides a composition comprising a composite of acitric acid polyester having the generic formula (A-B-C)_(n), wherein Ais a linear aliphatic dihydroxy monomer; B is citric acid, C is a linearaliphatic dihydroxy monomer, and n is an integer greater than 1; and abioceramic used for implantable tissue devices, wherein less than 75 wt.% weight ratio of said composition comprises said bioceramic. In otherembodiments, at least 30 wt. % weight ratio of said compositioncomprises said bioceramic. In specific and alternative embodiments, thebioceramic component forms about 5 wt. %, 10 wt. %, 15 wt. %, 20 wt. %,25 wt. %, 30 wt. %, 35 wt. %, 40 wt. %, 45 wt. %, 50 wt. %, 55 wt. %, 60wt. %, 65 wt. %, 70 wt. %, 75 wt. %, 80 wt. %, 85 wt. %, 90 wt. %, 95wt. %, or greater than 95 wt. % of the composition.

Preferred compositions are those in which A is a linear diol comprisingbetween about 2 and about 20 carbons. In other preferred compositions, Cis a linear diol comprising between about 2 and about 20 carbons. Instill other preferred compositions, both A and C are the same lineardiol. In alternative embodiments, A and C are different linear diols. Inspecific compositions, the linear diol is 1, 8-octanediol.

In exemplary embodiments, the linear aliphatic dihydroxy poly 1,8-octanediol co-citric acid. In still other exemplary embodiments, thelinear aliphatic dihydroxy poly 1, 10-decanediol co-citric acid.

The bioceramic may be any ceramic typically used in medicalapplications. For example, exemplary such bioceramics may be selectedfrom the group consisting of calcium phosphate bioceramics,alumina-based bioceramics; zirconia-based bioceramics; silica-basedbioceramics, and pyrolytic carbon-based bioceramics. Combinations ofbioceramics may be used. In certain preferred embodiments, thebioceramic is a calcium phosphate bioceramic at a weight percentage offrom 30 wt. % to about 75 wt. % of the total weight of the composition.In other embodiments, the bioceramic is hydroxyapatite (HA) at a weightpercentage of between about 40 wt. %±5 wt. % to about 70 wt. %±5 wt. %HA to 35 wt. %±5 wt. % to about 25 wt. %±5 wt. % citric acid polyester.

The compositions of the invention are such that they produce compositesof preferred bending strength. In preferred embodiments, the compositehas a bending strength of from about 33.9 to about 41.4 MPa. In stillother preferred embodiments, the composites have a preferred compressionstrength, wherein the compression strength is preferably from about 32to about 75 MPa. The composites alternatively or in addition may have adefined tensile strength in the range of from about 6 to about 10 MPa.The composites also may be characterized according to their shearstrength. The preferred shear strength of the composites is betweenabout 23 to about 28 MPa. The composites may be defined according totheir bending modulus. The bending modulus is preferably from about0.275 to about 0.502 GPa. In other embodiments, the composites arecharacterized by having a compression modulus of from about 0.19 toabout 0.45 GPa. In still other embodiments, the composites may becharacterized by a tensile modulus of from about 0.02 to about 0.34 GPa.The composites of the invention may advantageously be characterized byone or more of these features.

In addition, the invention contemplates composites of the inventionwhich in addition to the citric acid polyester and bioceramic furthercomprise a polymer selected from the group consisting ofpoly(hydroxyvalerate), poly(lactide-co-glycolide),poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate),polyorthoester, polyanhydride, poly(glycolic acid), poly(glycolide),poly(L-lactic acid), poly(L-lactide), poly(D,L-lactic acid),poly(D,L-lactide), poly(caprolactone), poly(trimethylene carbonate), andpolyester amide.

In specific embodiments, the composition is molded into an orthopedicfixation device. Preferred such devices include but are not limited tobone screws, bone pins, bone rods, and bone plates.

Also contemplated herein is an artificial bone, wherein said bone iscomprised of a composite of the present invention.

Another aspect of the invention describes a substrate for use in animplantable device comprising a composite of the invention prepared,molded or fabricated into an orthopedic fixation device or an artificialbone structure. In preferred embodiments, the substrate furthercomprises a surface modification to facilitate implantation of saiddevice with a decreased risk of implant rejection.

Also provided herein is a method of producing an implantable device,comprising: preparing a composition according to the present inventionand molding the composition into an orthopedic fixation device or anartificial bone for implantation. Such a method may be used in thepreparation of a bone screw, a bone pin, and a bone plate.

Also provided herein is an implantable device comprising a polymercomposition of the present invention.

Other features and advantages of the invention will become apparent fromthe following detailed description. It should be understood, however,that the detailed description and the specific examples, whileindicating preferred embodiments of the invention, are given by way ofillustration only, because various changes and modifications within thespirit and scope of the invention will become apparent to those skilledin the art from this detailed description.

BRIEF DESCRIPTION OF THE DRAWINGS

The following drawings form part of the present specification and areincluded to further illustrate aspects of the present invention. Theinvention may be better understood by reference to the drawings incombination with the detailed description of the specific embodimentspresented herein.

FIG. 1 Production of bone screws of POC-HA composite with 40 wt. % HAobtained by compression molding and machining methods. POC-HA compositeswere prepared by an in situ post-polymerization of HA and pre-POCblending at 80° C. for 3 days and 120° C. for 1 day under vacuum.

FIG. 2 SEM images of surface of POC-HA composites with a) 40 wt. % HAand b) 65 wt. % HA.

FIG. 3 Weight loss of POC-HA composites with HA fraction of 40, 50, 60,to 65 wt. % in vitro (PBS at 37° C.) at 2, 6, 12 and 20 weeks. POC-HAcomposites were prepared at 80° C. for 3 days and 120° C. under vacuumfor 1 day

FIG. 4 Mineralization in SBF for POC at a) 3 days and b) 15 days, forPOC-HA with 40 wt. % HA at c) 3 days, d) 15 days, and for POC-HA with 65wt. % HA e)3 days and f) 15 days. Magnification of all images: x 4.5K.

FIG. 5 LM images of POC seeded with HOB in vitro at a) 3 days b) 14days; SEM images of POC-HA composites seeded with human osteoblasts invitro for 40 wt. % HA at c) 3 days d) 15 days and for 65 wt. % at e) 3days and f) 14 days.

FIG. 6 is a schematic representation of the synthesis of poly(1,8-octanediol-co-citric acid).

DESCRIPTION OF THE PREFERRED EMBODIMENTS

In co-pending application 60/721,687 and PCT/US2004/030631, (eachincorporated herein by reference in its entirety) there is described thesynthesis and characterization of elastomeric and biodegradablepolyesters, referred to as poly (diol citrates (POC); seePCT/US2004/030631) and composites comprising POC with a second polymer(see U.S. 60/721,687 and applications depending therefrom). Themechanical properties and degradation rates of POC polymers can becontrolled with synthesis conditions of the polycondensation reactionand choice of diol, and their preparation does not involve any harshsolvents or exogenous catalysts [15, 16]. Furthermore, poly (diolcitrates) can be very inexpensive, relative to poly (α-hydroxy acid)biodegradable polymers.

In the present invention, it is shown that a composite of a poly (diolcitrate) with HA would have the desired characteristics of a bioceramicwith improved processability, mechanical properties, and degradationcharacteristics. Poly (1, 8-octanediol-co-citrate) (POC; the productionof which is shown in FIG. 6) was selected for the preparation ofexemplary composites of the present invention because of its fasterdegradation rate (a few months to 1 year) than PLLA (3-5 years) andbecause its mechanical properties can be tailored by simply changingreaction conditions such as reaction temperature and time, and the ratioof 1, 8-octanediol to citric acid [15, 16]. POC has also been shown tobe biocompatible and could potentially enhance the biointegration of thesurrounding soft tissue as in the case of fixation of a ligament graft[15, 16]. Moreover, these materials are inexpensive and easy tosynthesize, an additional advantage for clinical application.

In preferred embodiments composites are prepared from POC. Themethodology that was used to post-polymerize the materials is unique.Polycondensation of POC can be conducted under no vacuum, no catalyst,and low reaction temperature (under 100° C., such as 60° C., 80° C.,even as low as 37° C.). Catalyst and high temperature could also beapplied if needed.

The present invention shows that the POC-HA composites have the desiredmechanical, degradation, mineralization, and cell compatibilitycharacteristics to serve as compositions for implantable devices such asorthopedic fixation devices as well as to serve as compositions for usein the production of artificial bone structures. The feasibility offabricating (i.e., molding, or machining) composite bone screws ofPOC-HA by compression molding and machining also is shown herein.

Compositions of poly(diol citrates) comprise a citric acid polyesterhaving the generic formula (A-B-C)n, wherein A and C could be any of thediols or any combination of the diols; B could be citric acid, malicacid or their combinations. The diols include aliphatic diols, brancheddiol, cyclodiol, triol, heteroatomcontaining diol (such asN-methyldiethanolamine, MDEA) and macrodiol or their combinations. Anycomposites composed of any biodegradable elastomers (e.g. polydiol-citric acid, polyurethanes, polycaprolactone and copolymersthereof) and any bioceramic (such as HA, TCP, OCP and bioglass) withfraction from 0 to 100 wt. %, for example, 30 to 95 wt. % ceramics canbe applied to fabricate fixation devices (such as pins, wires, tacks,plates, rods, screws) for clinic and cell scaffolds for tissueengineering and drug delivery.

Orthopedic fixation devices such as bone screws are often used inorthopedic, craniofacial and oral-maxillofacial surgery. The vastmajority of these devices are made from metals, which can cause unwantedtissue reactions, and lead to significant bone removal if a secondaryintervention is required. Alternatively, biodegradable polymers such aspoly (L-lactide) (PLLA) have been used for the fabrication of somefixation devices where significant weight bearing is not an issue forthe proper function of the device. Unfortunately, these devices, inparticular PLLA bone screws, are not osteoconductive, have a slowdegradation rate (3-5 years), can fracture during the fixationprocedure, and are significantly more expensive than their metalcounterparts.

One way to deal with the osteointegration deficiencies of polymers hasbeen to blend them with bioceramics such as hydroxyapatite (HA) andtricalcium phosphate (TCP)[12-14, 22]. Bioceramics have been shown to beosteoconductive, but are brittle and hard to process into usefulfixation devices for orthopaedic applications. As a result, severalresearchers have developed and investigated composites of HA or TCP withpoly (a hydroxy acids) such as PLLA. Studies have found such compositesto osteointegrate more readily than the pure polymer, supporting thefurther study of HA/polymer composites. Nevertheless, the polymercomponent remains a relatively large percentage of the composites,typically 70 wt. %, and in the case of PLLA that is used commercially,the time to total degradation after its function has been completed isstill too long.

As shown in the Examples herein below, the novel bioceramic composites,based on poly (diol citrates), have enhanced osteointegration potentialrelative to current biodegradable fixation devices. In the presentinvention it is shown that it is possible to prepare bone screws andother orthopedic fixation devices that consist mostly of the bioceramiccomponent. This maximizes the osteointegration while employing adegradable and relatively inexpensive elastomer as the macrophasebinder. The inventors showed that poly (1,8-octanediol co-citrate)(POC), improves the processability and mechanical properties ofbioceramic bone screws due to its biocompatibility, mechanicalproperties, controllable degradation rates (a few months to 1 year), andmild synthesis conditions[1,5]. An important criterion for the POC-HAcomposites was the ability to process samples via molding and machiningmethods. The Examples below demonstrate the successful synthesis ofPOC-HA composites with HA compositions of 40, 50, 60, and 65 wt. % thatwere readily molded and machined to make bone screws. While HApercentages of 70 or higher may be fabricated, those composites are notreadily amenable to molding or machining and hence are less desirable,but may nonetheless be useful composites if molding is not required. AnHA content below 40 wt. % resulted in composites that were toonibber-like and difficult to machine. Such composites with an HA contentof less than 40 wt. % may nonetheless be useful in applications that donot require rigidity, e.g., tissue culture scaffolds and the like.

The mechanical property measurements of the POC-HA composites werewithin the range of values reported for biodegradable polymers andcomposites used or proposed for bone fixation devices [23]. Reportedmechanical properties for polymers and composites have included bending,compression, tensile strengths, and shear strengths, whose values rangedfrom 40-412 MPa, 78-130 MPa, 0.6-290 MPa and 19-250 MPa, respectively.Reported values for bending, compression, and tensile moduli ranged from1.6-124.4 GPa, 4.8-8.0 GPa, and 0.01-29.9 GPa, respectively[12, 23]. ThePOC-HA composites tested in this study had bending, compression,tensile, and shear strengths that ranged from 33.9-41.4 MPa, 32-75 MPa,6-10 MPa, and 23-28 MPa, respectively. Bending, compression, and tensilemoduli for POC-HA composites ranged from 0.275-0.502 GPa, 0.19-0.45 GPa,and 0.02-0.34 GPa, respectively. Except for the bending and compressionmoduli, the mechanical properties of POC-HA are comparable to those ofother biomaterials proposed for bone fixation.

The mechanical properties of the POC-HA composites were increased byincreasing the HA component. It is also possible to modulate themechanical properties with the reaction conditions, i.e., reactiontemperature and time, and choice of diol for the polycondensationreaction. Given the teachings of the present invention, the compositesof the invention can readily be adapted for in vivo use in the intendedapplication. The polymer-HA composites are expected to integrate withbone. It is contemplated that the mass percent and rate of degradationof the polymer component are parameters that may influence the functionand in vivo integration of the composite. POC samples that weresynthesized under the same conditions as the synthesis of the POC-HAcomposites lost 46 wt. % of their mass in 3 months. Further, POC whennot combined with HA has previously been shown to completely degradewithin 6 months when incubated in PBS at 37° C. [1,5]. Without beingbound by any theory or mechanism of action, it is possible that thelower degradation rates reported for the POC-HA composites may be due todiffering extents of the polycondensation reaction due to the lower masspercentages of polymer and the presence of thermally conductive HAparticles. Both of those parameters are expected to affect the degree ofcross-linking relative to pure POC for the same reaction temperature andtime. Furthermore, it is also possible for the POC to covalently reactwith OH groups on the HA particles effectively crosslinking the POC-HAmatrix [24-26]. HA would also serve as a buffer to the acidic functionalgroups and products generated from POC degradation, minimizing anyautohydrolytic effect on degradation. Further, the degradation of thePOC component can be significantly increased by “doping” with glycerolor N-methyldiethanolamine (MDEA)[16].

In vivo, HA has been shown to induce the deposition of calcium phosphatemineral on the surface of ceramic implants and bond to bone [27, 28].The capacity of POC and POC-HA to mineralize was assessed in vitro usinga modified simulated body fluid solution. Based on the SEM and EDXanalysis, POC-HA composites with 40-65 wt. % HA in SBF successfullyinduced surface mineralization. The mineralization process involved anucleation phase and a growth phase as evidenced by the completecoverage of the samples after 15 days of incubation in SBF [20].However, POC was not conducive to mineralization. The apatite or calciumphosphate mineral deposition [29] may contribute to improved bonebonding in vivo and help fill in any void volumes or pores left behindby degraded POC. Depending on which bioceramic is chosen for a finalapplication, most of the mass of the screw is expected to be integrated(when the bioceramic is an HA-type ceramic) or remodelled (when thebioceramic a TCP-type ceramic) by bone tissue and the remaining POCshould be totally degraded within 2 years of implantation.

POC has been shown to be compatible (i.e. as per cell adhesion,proliferation, and differentiation assays) with several cell typesincluding human and pig endothelial cells, human and pig smooth musclecells, bovine chondrocytes, and bovine fibroblasts [16, 30]. It was alsoshown to be biocompatible in vivo in a rat subcutaneous implantationmodel [1,5]. In the present invention, the favorable cell adhesion andspreading characteristics of POC and POC-HA composites were confirmed invitro with the use of primary human osteoblasts. The cells adhered andformed a confluent monolayer on all of the POC-HA composites evaluated(40, 50, 60, and 65 wt. % HA) (see e.g. FIG. 5) providing evidence thatthe composites of the present invention will be readily biocompatibilein vivo. Such determinations may further be corroborated using a bonedefect model.

From the above discussion, it is readily apparent that POC-HA compositescan be fabricated and molded into a variety of orthopedic fixationdevices. For example, POC-HA bone screws with an HA content of 65 wt. %were successfully prepared. It will readily be apparent that thesecompositions provide tremendous advantages over the existingtechnologies. Advantages of such POC-bioceramic composites include oneor more of the following: a) simple synthesis and in-situ crosslinkingpolymerization at relatively mild temperatures while avoiding the use ofexogenous catalysts and toxic solvents, b) incorporation of a highpercentage of the bioceramic component, potentially enhancingosteointegration, 3) a polymer component that should degrade completelywithin two years rather than three to five years as in the case withPLLA, and 4) decreased cost relative to the use of poly (α-hydroxylacids) such as PLLA. Moreover, the mechanical properties of POC-HAcomposites can be adjusted with the percent of HA in the composite andthe material's surface supported mineralization and osteoblast adhesionand proliferation. The bioceramic particle size can be readily adjustedand its effects on mechanical properties and in vivo bone integrationcharacteristics of the composite can be readily assessed.

In support of the above discussion, the following discussion provides afurther brief explanation of exemplary individual components of thecomposites. Poly(diol citrates) are a family of biodegradable andbiocompatible elastomers that have shown significant potential for softtissue engineering, see e.g., U.S. patent application see U.S.60/721,687 and applications depending therefrom. However, while thoseprior compositions are useful in the production of matrices for tissueculture and implantable tissue patches, those compositions are ofinsufficient rigidity to serve in orthopedic indications. It isdesirable to increase the strength and stiffness of those composites inorder to serve orthopedic purposes. The methods of the present inventionare directed to strengthening such POC based polymers. Methods andcompositions for preparing POC are described in detail inPCT/US2004/030631 and U.S. 60/721,687.

As noted herein, the POC is strengthened and stiffened for use inorthopedic applications by incorporating bioceramics into the POC-basedelastomeric matrix. “Bioceramics” materials typically are made ofinorganic salts that include the ions of calcium and phosphate, or inother examples include sulfate and carbonate. Bioceramics fulfill aunique function as biomedical materials and are used in a wide varietyof applications in the human body.

Preferably, the bioceramics useful in the invention are substantiallynon-toxic, biodegradable, bioerodable, and bioresorbable. The terms“biodegradable” and “bioerodable” as used herein similarly refer to amaterial property where biological, biochemical, metabolic processes,and the like may effect the erosion or degradation of the material overtime. Such degradation or erosion is due, at least in part, to contactwith substances found in the surrounding tissues, body fluids, andcells, or via cellular action, enzymatic action, hydrolytic processes,and other similar mechanisms in the body. The term “bioresorbable” asused herein refers to materials that are used by, resorbed into, or areotherwise eliminated from the body of the patient via existingbiochemical pathways and biological processes. For example, inembodiments where the bioceramic comprises calcium phosphate,bioresorbed calcium phosphate may be redeposited as bone mineral, beotherwise reutilized within the body, or be excreted. It is understoodthat some materials become bioresorbable following biodegradation orbioerosion of their original state, as described above. Preferably, thebiocompatible material is such that it does not elicit a substantialdetrimental response in the host, including but not limited to an immunereaction, such as an inflammatory response, tissue necrosis, and thelike that will have a negative effect on the patient. In the event thatsuch a negative effect may be seen, preferably, the material may betreated with a composition that allows the host to avoid such an adverseresponse to the material.

Bioceramic materials typically are made of salts of alumina; zirconia;calcium phosphates; silica based glasses or glass ceramics; or pyrolyticcarbons. The salts used to prepare the bioceramics and the bioceramicmatrices fabricated therefrom are commercially available or are readilyprepared via known procedures. Bioceramics include calcium salts ofcarbonate, sulfate, phosphate, and the like. Exemplary bioresorbablecalcium salts effective in the composition of this invention includecalcium carbonate, calcium sulfate, calcium sulfate hemihydrate, alsoknown as plaster of Paris, and certain porous or precipitated forms ofcalcium phosphate, and the like. The porous bioceramic matrix may alsobe fabricated from any number of natural bone sources, such as autograftor allograft material, or synthetic materials that are compositionallyrelated to natural bone.

Calcium phosphate ceramics are in general prepared by sintering moresoluble calcium salts, for example Ca(OH)₂, CaCO₃, and CaHPO₄, with aphosphorus-containing compound such as P₂O₅. Such preparations ofcalcium phosphate ceramics are known to those of skill in the art andhave been described e.g., in U.S. Pat. Nos. 3,787,900; 4,195,366;4,322,398; 4,373,217 and 4,330,514 (each incorporated herein byreference). Exemplary calcium phosphates for use in the inventioninclude, but are not limited to, calcium metaphosphate, dicalciumphosphate dihydrate, calcium hydrogen phosphate, tetracalcium phosphates(TCPs), heptacalcium decaphosphate, tricalcium phosphates, calciumpyrophosphate dihydrate, crystalline hydroxyapatite, poorly crystallineapatitic calcium phosphate, calcium pyrophosphate, monetite, octacalciumphosphate, and amorphous calcium phosphate.

Chemical formulae for calcium phosphate ceramics also are provided in arange of crystalline morphologies, all of which may be used infabricating the bioceramic matrix, as described by U.S. Pat. Nos.6,331,312 and 6,027,742. Such calcium phosphates have been described aspoorly-crystalline calcium phosphate (PCA) with an apatitic structure.Other examples include tricalcium phosphate, tetracalcium phosphate andother mixed-phase or polycrystalline calcium phosphate materialsreported in U.S. Pat. Nos. 4,880,610 and 5,053,312 to Constanz et al.,the disclosures of which are incorporated herein by reference.

Particularly preferred bioceramics for use in the present inventioninclude calcium phosphate apatites, such as hydroxyapatite (HA,Ca₁₀(PO₄)₆(OH)₂) described by R. E. Luedemann et al., Second WorldCongress on Biomaterials (SWCB), Washington, D.C., 1984, p. 224,fluoroapatites, tricalciumphosphates (TCP), such as Synthograft,dicalciumphosphates (DCP), and mixtures of HA and TCP, as described byE. Gruendel et al., ECB, Bologna, Italy, 1986, Abstracts, p. 5, p. 32);mixed-metal salts such as magnesium calcium phosphates, and beta-TCMP,as described by A. Ruggeri et al., Europ. Congr. on Biomaterials (ECB),Bologna, Italy, 1986, Abstracts, p. 86; aluminum oxide ceramics;bioglasses such as SiO₂—CaO—Na₂O—P₂O₅, e.g. Bioglass 45S (SiO₂ 45 wt %.CaO 24.5%, Na₂O 24.5% and P₂O.sub.5 6%) described by C. S. Kucheria etal., SWBC, Washington, D.C., 1984, p. 214, and glass ceramics withapatites (MgO 4.6 wt %, CaO 44.9%, SiO₂ 34.2%, P₂O₅ 16.3% and CaF 0.5%)described by T. Kokubo et al., SWBC, Washington, D.C., 1984, p. 351;bioceramics incorporating organic ions, such as citrate, as described inU.S. Pat. No. 5,149,368 to Liu et al.; and commercial materials, such asDurapatite, Calcitite, Alveograf, and Permagraft; the disclosures ofwhich are incorporated herein by reference.

In addition to the calcium-based bioceramics, it is contemplated thatbioactive glass compositions may also be used. Such bioceramics includeSiO₂, Na₂O, CaO, P₂O₅, Al₂O₃, and CaF₂. It is appreciated that theabove-described calcium salts may be used alone or may be mixed toprepare the bioceramics described herein. Alumina and Zirconia are knownfor their general chemical inertness and hardness. These properties areexploited for implant purposes, where it is used as an articulatingsurface in hip and knee joints. The ability of these materials to bepolished to a high surface finish makes them ideal candidates for thiswear application. Porous alumina has also been used as a bone spacer,where sections of bone have had to be removed due to disease. In thisapplication, it acts as a scaffold for bone ingrowth. Single crystalalumina or sapphire has also been used.

Pyrolytic carbon is a bioceramic commonly used in artificial heartvalves and has been the most popular material for this application forthe last 30 years. Properties that make this material suitable for thisapplication include good strength, wear, resistance and durability, andmost importantly, thromboresistance, or the ability to resist bloodclotting. Pyrolytic carbon is also used for small orthopaedic jointssuch as fingers and spinal inserts.

The POC-bioceramic materials may be prepared as porous or channeledstructures. It is understood that the nature and size of these pores orchannels may affect bioresorption. In certain aspects, the pores of thestructure are interconnected forming an open-cell porous structure. Itis understood that each of the foregoing materials may possess differingbioresorption characteristics obtainable in the treatment subject andsuch characteristics may be advantageously chosen via routineexperimentation for particular variations of the processes and methodsdescribed herein. It is also understood that both chemical compositionand crystal morphology may affect bioresorption rates. For example,bioceramics fabricated from mixtures of calcium phosphate and calciumcarbonate or calcium phosphate and calcium sulfate typically undergoresorption at higher rates than bioceramics fabricated from calciumphosphate alone. Furthermore, highly crystalline bioceramics typicallyundergo resorption at rates slower than poorly crystalline or amorphousbioceramics.

The porous microstructure of the bioceramics may be achieved by heatconsolidation or sintering of bioceramic powders in appropriate molds.The porous matrices may be macroporous or microporous. Microporousmatrices typically have pores in the range from about 1 to about 100microns in size, while macroporous matrices typically have pores in therange from about 100 to about 1000 microns in size. In certainembodiments the pore size in a given range is substantially uniform. Thepores in the matrix account for the void volume thereof. Such voidvolume may be from about 30% to about 80%, and illustratively about 50%to about 70% of the matrix volume. The pores are typicallyinterconnecting, and in some cases to a substantial degree. The poresmay form an open-cell configuration in some embodiments. In embodimentswhere the void volume constitutes a substantial portion of the matrixvolume, the pores are typically close together. Illustratively, adjacentpores are separated by less than 100 microns, and in other embodimentsseparated by less than about the average of the diameters of theadjacent pores.

In embodiments where the POC-bioceramic composites are used for thepreparation of bone-replacement materials, the pores may be arranged inpredetermined patterns that correspond to bone-healing orbone-remodeling patterns, Haversian systems, and othernaturally-occurring patterns in bone. Commercially available bioceramicmatrices include e.g., Pro Osteon 200 and Pro Osteon 500 (hydroxyapatitebone-graft substitutes having interconnected porous structures with poresizes of 200 or 500 microns, similar to that of cancerous bone)available from Interpore International, Irvine, Calif.; Vitoss Blocks(calcium phosphate porous structure having ca. 90% porosity, with poresizes from 1 to 1000 microns in diameter) available from Orthovita Inc.,Malvern, Pa.; and synthetic porous hydroxyapatite (made by a patentedfoam process having controlled porosity and pore sizes) available fromHi-Por Ceramics, United Kingdom. Such compositions may readily be usedin the present invention.

While the present invention is directed to production of components fororthopedic devices, it is contemplated that in addition to thebioceramic, the POC composites also may be reinforced with a secondmaterial. For example, the POC-HA composites also may comprise a secondbiodegradable and biocompatible polymer. Two such polymer arepoly(L-lactic acid) (PLLA) and poly(lactic-co-glycolic acid) (PLGA).These polymers are rigid and strong and have been used in many tissueengineering applications. Furthermore, the rate of degradation could betailored to match that of the surrounding elastomeric matrix. Poly(L-lactic acid) has a degradation time of greater than two years whilepoly(glycolic acid) has a degradation time of 1-2 months. By changingthe ratio of lactic to glycolic acid, the degradation rate could bevaried from fast (1-2 months) to slow (>2 years). For tissueengineering, the rate of degradation of the polymer scaffold shouldmatch that of tissue regrowth.

The reinforcing polymer may be a biodegradable polymer or anon-biodegradable polymer, but preferably is a biodegradable polymer.Biodegradable polymers include, but are not limited to collagen,elastin, hyaluronic acid and derivatives, sodium alginate andderivatives, chitosan and derivatives gelatin, starch, cellulosepolymers (for example methylcellulose, hydroxypropylcellulose,hydroxypropylmethylcellulose, carboxymethylcellulose, cellulose acetatephthalate, cellulose acetate succinate, hydroxypropylmethylcellulosephthalate), casein, dextran and derivatives, polysaccharides,poly(caprolactone), fibrinogen, poly(hydroxyl acids), poly(L-lactide)poly(D,L lactide), poly(D,L-lactide-co-glycolide),poly(L-lactide-co-glycolide), copolymers of lactic acid and glycolicacid, copolymers of .ε-caprolactone and lactide, copolymers of glycolideand .ε-caprolactone, copolymers of lactide and 1,4-dioxane-2-one,polymers and copolymers that include one or more of the residue units ofthe monomers D-lactide, L-lactide, D,L-lactide, glycolide,ε-caprolactone, trimethylene carbonate, 1,4-dioxane-2-one or1,5-dioxepan-2-one, poly(glycolide), poly(hydroxybutyrate),poly(alkylcarbonate) and poly(orthoesters), polyesters,poly(hydroxyvaleric acid), polydioxanone, poly(ethylene terephthalate),poly(malic acid), poly(tartronic acid), polyanhydrides,polyphosphazenes, poly(amino acids). The biodegradable polymers usedherein may be copolymers of the above polymers as well as blends andcombinations of the above polymers. (see generally, Illum, L., Davids,S. S. (eds.) “Polymers in Controlled Drug Delivery” Wright, Bristol,1987; Arshady, J. Controlled Release 17:1-22, 1991; Pitt, Int. J. Phar.59:173-196, 1990; Holland et al., J. Controlled Release 4:155-0180,1986).

In particular preferred embodiments, the biodegradable or resorbablepolymer is one that is formed from one or more monomers selected fromthe group consisting of lactide, glycolide, ε-caprolactone, trimethylenecarbonate, 1,4-dioxan-2-one, 1,5-dioxepan-2-one, 1,4-dioxepan-2-one,hydroxyvalerate, and hydroxybutyrate. In one aspect, the polymer mayinclude, for example, a copolymer of a lactide and a glycolide. Inanother aspect, the polymer includes a poly(caprolactone). In yetanother aspect, the polymer includes a poly(lactic acid),poly(L-lactide)/poly(D,- L-Lactide) blends or copolymers of L-lactideand D,L-lactide. In yet another aspect, the polymer includes a copolymerof lactide and ε-caprolactone. In yet another aspect, the polymerincludes a polyester (e.g., a poly(lactide-co-glycolide). Thepoly(lactide-co-glycolide) may have a lactide:glycolide ratio rangesfrom about 20:80 to about 2:98, a lactide:glycolide ratio of about10:90, or a lactide:glycolide ratio of about 5:95. In one aspect, thepoly(lactide-co-glycolide) is poly(L-lactide-co-glycolide; see e.g.,U.S. Pat. No. 6,531,146 and U.S. application No. 2004/0137033.). Otherexamples of biodegradable materials include polyglactin, andpolyglycolic acid.

Representative examples of non-biodegradable compositions includeethylene-co-vinyl acetate copolymers, acrylic-based andmethacrylic-based polymers (e.g., poly(acrylic acid), poly(methylacrylicacid), poly(methylmethacrylate), poly(hydroxyethyl methacrylate),poly(alkylcynoacrylate), poly(alkyl acrylates), poly(alkylmethacrylates)), polyolefins such as poly(ethylene) or poly(propylene),polyamides (e.g., nylon 6,6), poly(urethanes) (e.g., poly(esterurethanes), poly(ether urethanes), poly(carbonate urethanes),poly(ester-urea)), polyesters (e.g., PET, polybutyleneterephthalate, andpolyhexyleneterephthalate), olyethers (poly(ethylene oxide),poly(propylene oxide), poly(ethylene oxide)-poly(propylene oxide)copolymers, diblock and triblock copolymers, poly(tetramethyleneglycol)), silicone containing polymers and vinyl-based polymers(polyvinylpyrrolidone, poly(vinyl alcohol), poly(vinyl acetatephthalate), poly(styrene-co-isobutylene-co-styrene), fluorine containingpolymers (fluoropolymers) such as fluorinated ethylene propylene (FEP)or polytetrafluoroethylene (e.g., expanded PTFE).

The polymers may be combinations of biodegradable and non-degradablepolymers. Further examples of polymers that may be used are eitheranionic (e.g., alginate, carrageenin, hyaluronic acid, dextran sulfate,chondroitin sulfate, carboxymethyl dextran, caboxymethyl cellulose andpoly(acrylic acid)), or cationic (e.g., chitosan, poly-1-lysine,polyethylenimine, and poly(allyl amine)) (see generally, Dunn et al., J.Applied Polymer Sci. 50:353, 1993; Cascone et al., J. Materials Sci.:Materials in Medicine 5:770, 1994; Shiraishi et al., Biol. Pharm. Bull.16:1164, 1993; Thacharodi and Rao, Int'l J. Pharm. 120:115, 1995;Miyazaki et al., Int'l J. Pharm. 118:257, 1995). Preferred polymers(including copolymers and blends of these polymers) includepoly(ethylene-co-vinyl acetate), poly(carbonate urethanes),poly(hydroxyl acids) (e.g., poly(D,L-lactic acid) oligomers andpolymers, poly(L-lactic acid) oligomers and polymers, poly(D-lacticacid) oligomers and polymers, poly(glycolic acid), copolymers of lacticacid and glycolic acid, copolymers of lactide and glycolide,poly(caprolactone), copolymers of lactide or glycolide andε-caprolactone), poly(valerolactone), poly(anhydrides), copolymersprepared from caprolactone and/or lactide and/or glycolide and/orpolyethylene glycol. Methods for making POC-PLLA or PLGA or other likecomposites are described in U.S. 60/721,687.

The composites of the invention are load bearing and may bear loadssimilar in magnitude to that borne by the tissue surrounding the defect,such as a bone structure of similar dimensions, or a bone structureconsisting primarily of cortical bone. The structures described hereinmay also possess mechanical properties similar to that of natural bone,or in particular cortical bone. These mechanical properties include, butare not limited to, tensile strength, impact resistance, Young'smodulus, compression strength, sheer strength, stiffness, and the like.It is appreciated that structures described herein possessing mechanicalproperties similar to those exhibited by the tissue surrounding suchimplanted structures may favorably influence the stress-shieldingeffect.

While it is appreciated that the above-described composites of theinvention will be fashioned into replacement bone or orthopedic fixationdevices, it is also contemplated that the devices may be used as drugdelivery system. Either the polymer, the bioceramic, or both may includea biologically-active agent, either singly or in combination, such thatthe composite structure or implant will provide a delivery system forthe agent at the site at which it is implanted. Thus, the agent mayadvantageously be delivered to adjacent tissues or tissues proximal tothe implant site. Biologically-active agents which may be used alone orin combination in the implant precursor and implant include, forexample, a medicament, drug, or other suitable biologically-,physiologically-, or pharmaceutically-active substance which is capableof providing local or systemic biological, physiological, or therapeuticeffect in the body of the patient. The biologically-active agent iscapable of being released from the solid implanted matrix into adjacentor surrounding tissue fluids during biodegradation, bioerosion, orbioresorption of the fixation device or artificial bone made from thecomposites of the invention.

Incorporation of bioceramic composites of the invention supportsmineralization and osteoblast adhesion and proliferation, and canpotentially enhance osteointegration. In vivo bone integrationcharacteristics of the biodegradable elastomer-bioceramic composites canbe adjusted with different kinds and sizes of bioceramic. Where thecomposites are used in facilitating bone repair, the composites mayadvantageously be impregnated with an “osteogenic agent” i.e., one whichpromotes, induces, stimulates, generates, or otherwise effects theproduction of bone or the repair of bone. The presence of an osteogenicagent in the site at which the composite is placed may elicit an effecton the repair of the defect in terms of shortening the time required torepair the bone, by improving the overall quality of the repair, wheresuch a repair is improved over situations in which such osteogenicagents are omitted, or may achieve contemporaneously both shortenedrepair times and improved bone quality. It is appreciated thatosteogenic agents may effect bone production or repair by exploitingendogenous systems, such as by the inhibition of bone resorption.

Thus, osteogenic agents in the composites of the invention may be usedto effect repair of the bone by stabilizing the defect to promotehealing thereby increasing healing rate, producing a more rapid new boneingrowth, and improving overall repair of the bone. The osteogenicagents may be synthetic molecules, drugs, or pharmaceuticals involvedin, or important to, bone biology, including statins, such aslovastatin, simvastatin, atorvastatin, and the like, fluprostenol,vitamin D, estrogen, a selective estrogen receptor modifier, or aprostaglandin, such as PGE-2. Growth factors or other proteins,peptides, receptor ligands, peptide hormones, lipids, or carbohydratesinvolved in, or important to, bone physiology may be used, including thebone morphogenic or bone morphogenetic proteins (BMPs), such as BMP-2,BMP-7, and BMP-9, chrysalin, osteogenic growth peptide (OGP), bone cellstimulating factor (BCSF), KRX-167, NAP-52, gastric decapeptide,parathyroid hormone (PTH), a fragment of parathyroid hormone,osteopontin, osteocalcin, a fibroblast growth factor (FGF), such asbasic fibroblast growth factor (bFGF) and FGF-1, osteoprotegerin ligand(OPGL), platelet-derived growth factor (PDGF), an insulin-like growthfactor (IGF), such as IGF-1 and IGF-2, vascular endothelial growthfactor (VEGF), transforming growth factor (TGF), such as TGF-alpha andTGF-beta, epidermal growth factor (EGF), growth and differentiationfactor (GDF), such as GDF-5, GDF-6, and GDF-7, thyroid-derivedchondrocyte stimulation factor (TDCSF), vitronectin, laminin,amelogenin, amelin, fragments of enamel, or dentin extracts, bonesialoprotein, and analogs and derivatives thereof.

In another embodiment the osteogenic agent is a cell or population ofcells involved in, or important to, bone biology, such as pluripotentstem cells, autologous, allogenic, or xenogeneic progenitor cells,chondrocytes, adipose-derived stem cells, bone marrow cells, mesenchymalstem cells, homogenized or comminuted tissue transplants, geneticallytransformed cells, and the like. Bone powders, including demineralizedbone powders and bone matrix, may also be used. Combinations of suchcell populations providing the osteogenic agent are also contemplatedherein.

The osteogenic agent may be present in the structure within the rangefrom about 0.1% to about 30% by weight, preferably in the range fromabout 1% to 9% by weight.

Other agents also may be used in the composites. It is contemplated thatsuch additives may serve to reduce barriers to repair and thus maximizethe potential of the osteogenic agent. Preferably, such agents arecapable of preventing infection in the host, either systemically orlocally at the defect site, are contemplated as illustrative usefuladditives. These additives include anti-inflammatory agents, such ashydrocortisone, dexamethasone, prednisone, and the like, NSAIDS, such asacetaminophen, salicylic acid, ibuprofen, and the like, selective COX-2enzyme inhibitors, antibacterial agents, such as penicillin,erythromycin, polymyxin B, viomycin, chloromycetin, streptomycins,cefazolin, ampicillin, azactam, tobramycin, cephalosporins, bacitracin,tetracycline, doxycycline, gentamycin, quinolines, neomycin,clindamycin, kanamycin, metronidazole, and the like, antiparasiticagents such as quinacrine, chloroquine, vidarabine, and the like,antifungal agents such as nystatin, and the like, antiviricides,particularly those effective against HIV and hepatitis, and antiviralagents such as acyclovir, ribarivin, interferons, and the like. Systemicanalgesic agents such as salicylic acid, acetaminophen, ibuprofen,naproxen, piroxicam, flurbiprofen, morphine, and the like, and localanaesthetics such as cocaine, lidocaine, bupivacaine, xylocalne,benzocaine, and the like, also can be used as additives in thecomposites.

The composites of the present invention are used to fabricate structuresfor use in orthopedic applications. Such structures preferably are bonefixation devices, e.g., bone screws, pins, plates and the like.Alternatively, the structures are those that can be used for repairingbone voids, fractures, non-union fractures, periodontal defects,maxillofacial defects, arthrodesis, and the like. The composites may befabricated into useful structures using e.g., compression molding andmachining methods can be applied to fabricate any desirable shapes offixation devices and scaffolds for orthopaedic surgery and tissueengineering. The mechanical properties and degradation of thebiodegradable elastomer bioceramic composites can be adjusted with thepercent and particle size (from micron to nanometer) of any bioceramics(HA, TCP, OCP and bioglass) in the composite besides ratio of diol andcitric acid and post-polymerization conditions such as temperature,vacuum, and polymerization time.

In addition, structures described herein may be used as reinforcement ofbone fractures, dental implants, bone implants, bone prostheses and thelike. It is appreciated that structures described herein may also begenerally used in conjunction with other traditional fixation,immobilization, and prosthetic methods. Fractures that may be treated bystructures made from the composites of the invention include fracturesof the proximal humerus, diaphyseal humerus, diaphyseal femur,trochanteric femur, and trochanteric humerus. In addition, thestructures may be used in the repair of osteoporosis-induced fractures,including those that involve a crushing-type injury, such as vertebralfractures, and the like. In such fractures, the porous osteoporotic bonecollapses into itself typically causing a void or bone defect at thesite of the fracture, in order to achieve secure stabilization of thefracture.

The various structures that can be prepared using the composites of theinvention may be fabricated by using methods known to those of skill inthe art. Typically, the bone fixation devices or other structures foruse in the orthopedic applications described herein may be fabricated bycompression molding. Compression molding processes include transfermolding and squeeze-flow molding.

In exemplary embodiments, compression molding is used. A composite ofthe invention, in a machined-block form, is placed on top of abioceramic matrix in a mold cavity. The mold is then heated to atemperature at about or above the melting temperature of the polymer.Minimal loading occurs during the heating step. Pressurization of themold is initiated once the molten polymer is fluid enough for diffusionthrough the porous structure. In addition, vacuum may be optionallyapplied during this process to prevent degradation or hydrolysis ofbiocompatible polymer. It is appreciated that applying a vacuum may alsofacilitate the diffusion of polymer into the matrix. Exemplary moldingof the composites into bone screws is further described in the examplesherein below.

In other embodiments, transfer molding is used. A composite of theinvention, in a machined-block form, is preheated to a temperature atabout or above the melting temperature of the polymer and subsequentlytransferred to a preheated mold cavity containing a porous bioceramicmatrix. Once the molten polymer is positioned, squeeze molding isinitiated by applying a load to a plunger, thereby pressurizing the moldcavity.

In still other embodiments, flow molding is used. A porous bioceramicmatrix having a small-diameter open core is used. In addition, thematrix has interconnected channels that are also connected to the opencore. The channels are arranged in a substantially radial pattern whenviewed in a given cross section of the matrix. The porous matrix isplaced in a mold cavity and the POC material is disposed into the opencore by either of the above-described methods of compression molding ortransfer molding. In either case this process allows orientation of thepolymer from in the matrix. Such orientation may further reinforce andfavorably influence the mechanical properties of the structuresdescribed herein.

Once the POC is disposed in the porous matrix by any of the methodsdescribed herein, including compression molding, transfer molding,squeeze-flow molding, and in-situ polymerization, the polymer may beoptionally crosslinked. Cross-linking may be accomplished by any of thevariety of known methods, including treatment with heat or irradiation,such as X-ray radiation, gamma irradiation, electron beam radiation, andthe like.

It should be understood that the composites of the invention may beprovided to a practitioner as bulk material that may be shaped by themedical practitioner on site. Alternatively, various prefabricatedshapes ready or near ready for implantation may be produced from thecomposites. Such bulk material may in the form of bars, blocks, billets,sheets, and the like. Such shapes include plates, plugs, cubes,cylinders, pins, tubes, chutes, rods, screws, including the screwsdescribed in U.S. Pat. No. 6,162,225 (bone screw fabricated fromallograft bone) the disclosure of which is incorporated herein byreference, and the like. In addition, shapes that tend to mimic theoverall dimensions of the bone may be made from the composites of theinvention. Shapes that tend to mimic the overall dimensions of the boneare particularly useful in the repair of fractures at risk of non-union.Such bulk shapes or particularly-dimensioned shapes may be obtained byemploying mold cavities possessing such dimensions. Alternatively, theparticularly-dimensioned shapes may be fabricated by machining the bulkstock.

EXAMPLE 1 Biodegradable Elastomeric Polymers

The compositions of the invention are based on biodegradable elastomericpolymers of poly(diol) citrate molecules. Such molecules typicallycomprising a polyester network of citric acid copolymerized with alinear aliphatic di-OH monomer in which the number of carbon atomsranges from 2 to 20. Polymer synthesis conditions for the preparation ofthese molecules vary from mild conditions, even at low temperature (lessthan 100° C.) and no vacuum, to tough conditions (high temperature andhigh vacuum) according the requirements for the materials properties. Bychanging the synthesis conditions (including, but not limited to,post-polymerization temperature, time, vacuum, the initial monomer molarratio, and the di-OH monomer chain length) the mechanical properties ofthe polymer can be modulated over a wide range. This series of polymersexhibit a soft, tough, biodegradable, hydrophilic properties andexcellent biocompatibility in vitro.

The poly(diol)citrate polymers used herein have a general structure of:(A-B-C)_(n)where A is a linear, aliphatic diol and C also is a linear aliphaticdiol. B is citric acid. The citric acid co-polymers of the presentinvention are made up of multiples of the above formula, as defined bythe integer n, which may be any integer greater than 1. It iscontemplated that n may range from 1 to about 1000 or more. It isparticularly contemplated that n may be 1, 2, 3, 4, 5, 6, 7, 8, 9, 10,11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28,29, 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 46,47, 48, 49, 50, or more. In preferred embodiments, the compositions ofpoly (diol citrates) that are used to prepare the implantable medicaldevice comprise a citric acid polyester having the generic formula(A-B-C)n, wherein A and C could be any of the diols or any combinationof the diols; B could be citric acid, malic acid or their combinations.The diols include linear or noon-linear aliphatic diols, branched diol,cyclodiol, triol, heteroatom containing diol (such asN-methyldiethanolamine, MDEA) and macrodiol or their combinations. Anymedical device coated with any biodegradable elastomers (e.g., polydiol-citric acid, polyurethanes, polycaprolactone and copolymers thereofis contemplated to be within the aspects of the present invention.

In preferred embodiments, the identity of “A” above is poly 1,10-decanediol and in another preferred embodiment the identity of A is1, 8-octanediol. However, it should be understood that this is merely anexemplary linear, aliphatic diol. Those of skill in the art are aware ofother aliphatic alcohols that will be useful in polycondensationreactions to produce poly citric acid polymers. Exemplary such aliphaticdiols include any diols of between about 2 carbons and about 20 carbons.While the diols are preferably aliphatic, linear, unsaturated diols,with the hydroxyl moiety being present at the C₁ and C_(x) position(where x is the terminal carbon of the diol), it is contemplated thatthe diol may be an unsaturated diol in which the aliphatic chaincontains one or more double bonds. The preferred identity for “C” in oneembodiment is 1, 8, octanediol, however as with moiety “A,” “C” may beany other aliphatic alcohols. While in specific embodiments, both A andC are both the same diol, e.g., 1, 8-octanediol, it should be understoodthat A and C may have different carbon lengths. For example, A may be 2,3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20 or morecarbons in length, and C may independently be 2, 3, 4, 5, 6, 7, 8, 9,10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20 or more carbons in length.Exemplary methods for the polycondensation of the citric acid with thelinear diols are provided in this Example. These materials are then usedas starting materials for the composites described in Example 2.

Synthesis of Poly (1,10-decanediol-co-citric acid) (PDC) In a typicalexperiment, 19.212 g citric acid and 17.428 g 1,10-decanediol were addedto a 250 ml three-neck round-bottom flask, fitted with an inlet adapterand an outlet adapter. The mixture was melted within 15 min by stirringat 160-165° C. in silicon oil bath, and then the temperature of thesystem was lowered to 120° C. The mixture was stirred for half an hourat 120° C. to get the pre-polymer. Nitrogen was vented throughout theabove procedures. The pre-polymer was post-polymerized at 60° C., 80° C.or 120° C. with and without vacuum for predetermined time from one dayto 3 weeks depending on the temperature to achieve the Poly(1,10-decanediol-co-citric acid). Nitrogen was introduced into thereaction system before the polymer was taken out from the reactionsystem;

Preparation of Poly(1,8-Octanediol-co-citric acid) (POC) In a typicalexperiment, 19.212 g citric acid and 14.623 g 1,8-octanediol were addedto a 250 mL three-neck round-bottom flask, fitted with an inlet adapterand an outlet adapter. The mixture was melted within 15 min by stirringat 160-165° C. in silicon oil bath, and then the temperature of thesystem was lowered to 140° C. The mixture was stirred for another 1 hrat 140° C. to get the pre-polymer. Nitrogen was vented throughout theabove procedures. The pre-polymer was post-polymerized at 60° C., 80° C.or 120° C. with and without vacuum for predetermined time (from one dayto 3 weeks depending on the temperature, with the lower temperaturesrequiring longer times) to achieve the Poly (1,8-octanediol-co-citricacid). Nitrogen was introduced into the reaction system before thepolymer was taken out from the reaction system.

Porous scaffolds of POC (tubular and flat sheets) were prepared via asalt leaching technique as follows: POC pre-polymer was dissolved intodioxane to form 25 wt % solution, and then the sieved salt (90-120microns) was added into pre-polymer solution to serve as a porogen. Theresulting slurry was cast into a poly(tetrafluoroethylene) (PTFE) mold(square and tubular shape). After solvent evaporation for 72 h, the moldwas transferred into a vacuum oven for post-polymerization. The salt inthe resulting composite was leached out by successive incubations inwater (produced by Milli-Q water purification system every 12 h for atotal 96 h. The resulting porous scaffold was air-dried for 24 hr andthen vacuum dried for another 24 hrs. The resulting scaffold was storedin a desiccator under vacuum before use. Porous scaffolds are typicallypreferred when cells are expected to migrate through a 3-dimensionalspace in order to create a tissue slice. Solid films would be used whena homogenous surface or substrate for cell growth is required such as anendothelial cell monolayer within the lumen of a vascular graft.

Using similar techniques porous scaffold of PDC or other poly(diol)citrates can be prepared. In other embodiments, biphasic scaffoldscan be prepared. Biphasic scaffolds consist of an outside porous phaseand an inside non-porous phase as depicted in the schematic drawingshown in FIG. 15 of PCT PCT/US2004/030631, incorporated herein byreference. The non-porous phase is expected to provide a continuoussurface for EC adhesion and spreading, mechanical strength, andelasticity to the scaffold. The porous phase will facilitate the 3-Dgrowth of smooth muscle cells. Biphasic scaffolds were fabricated viafollowing procedures. Briefly, glass rods (˜3 mm diameter) were coatedwith the pre-polymer solution and air dried to allow for solventevaporation. Wall thickness of the tubes can be controlled by the numberof coatings and the percent pre-polymer in the solution. The pre-coatedpre-polymer was partially post-polymerized under 60° C. for 24 hr; thepre-polymer-coated glass rod is then inserted concentrically in atubular mold that contains a salt/pre-polymer slurry. Thepre-polymer/outer-mold/glass rod system is then placed in an oven forfurther post-polymerization. After salt-leaching, the biphasic scaffoldwas then de-molded from the glass rod and freeze dried. The resultingbiphasic scaffold was stored in a desiccator before use. The samematerials or different materials from the above family of elastomers canbe utilized for both phases of the scaffold. Other biomedical materialswidely used in current research and clinical application such aspolylactide (PLA), polycaptrolactone (PCL), poly(lactide-co-glycolide)(PLGA) may also be utilized for this novel scaffold design.

The thickness, degradation, and mechanical properties of the insidenon-porous phase can be well controlled by choosing various pre-polymersof this family of elastomers, pre-polymer concentration, coating timesand post-polymerization conditions (burst pressure can be as high as2800 mmHg). The degradable porous phase and non-porous phases areintegrated since they are formed in-situ via post-polymerization. Thecell culture experiments shown in FIG. 16 confirm that both HAEC andHASMC can attach and grow well in biphasic scaffolds. The resultssuggest that a biphasic scaffold design based on poly(diol citrate) is aviable strategy towards the engineering of small diameter blood vessels.

Synthesis of Poly (1,6-hexanediol-co-citric acid) (PHC). In a typicalexperiment, 19.212 g citric acid and 11.817 g 1,6-hexanediol were addedto a 250 ml three-neck round-bottom flask, fitted with an inlet adapterand an outlet adapter. The mixture was melted within 15 min by stirringat 160-165° C. in a silicon oil bath, and then the temperature of thesystem was lowered to 120° C. The mixture was stirred for half an hourat 120° C. to get the pre-polymer. Nitrogen was vented throughout theabove procedures. The pre-polymer was post-polymerized at 60° C., 80° C.or 120° C. with and without vacuum for a predetermined time from one dayto 3 weeks, depending on the temperature, to achieve the Poly(1,6-hexanediol-co-citric acid). Nitrogen was introduced into thereaction system before the polymer was taken out from the reactionsystem.

Synthesis of Poly (1,12-dodecanediol-co-citric acid) PDDC. In a typicalexperiment, 19.212 g citric acid and 20.234 g 1,12-dodecanediol wereadded to a 250 ml three-neck round-bottom flask, fitted with an inletadapter and an outlet adapter. The mixture was melted within 15 min bystirring at 160-165° C. in silicon oil bath, and then the temperature ofthe system was lowered to 120° C. The mixture was stirred for half anhour at 120° C. to get the pre-polymer. Nitrogen was vented throughoutthe above procedures. The pre-polymer was post-polymerized at 60° C.,80° C. or 120° C. with and without vacuum for predetermined time fromone day to 3 weeks depending on the temperature to achieve the Poly(1,12-dodecanediol-co-citric acid). Nitrogen was introduced into thereaction system before the polymer was taken out from the reactionsystem.

Synthesis of Poly(9,8-octanediol-co-citric acid-co-glycerol) In atypical experiment (Poly(1,8-octanediol-co-citric acid-co-1% glycerol),23.0544 g citric acid, 16.5154 g 1,8-octanediol and 0.2167 g glycerolwere added to a 250 ml three-neck round-bottom flask, fitted with aninlet adapter and an outlet adapter. The mixture was melted within 15min by stirring at 160-165° C. in silicon oil bath, and then thetemperature of the system was lowered to 120° C. The mixture was stirredfor another hour at 140° C. to get the pre-polymer. Nitrogen was ventedthroughout the above procedures. The pre-polymer was post-polymerized at60° C., 80° C. or 120° C. with and without vacuum for predetermined timefrom one day to 3 weeks depending on the temperature to achieve the Poly(1,8-octanediol-co-citric acid-co-1% glycerol). Nitrogen was introducedinto the reaction system before the polymer was taken out from thereaction system.

Synthesis of Poly(1,8-octanediol-citric acid-co-polyethylene oxide). Ina typical experiment, 38.424 g citric acid, 14.623 g 1,8-octanediol and40 g polyethylene oxide with molecular weight 400 (PEO400)(100 g PEO1000and 200 g PEO2000 respectively) (molar ratio: citricacid/1,8-octanediol/PEO400=1/0.5/0.5) were added to a 250 ml or 500 mlthree-neck round-bottom flask, fitted with an inlet adapter and anoutlet adapter. The mixture was melted within 15 min by stirring at160-165° C. in silicon oil bath, and then the temperature of the systemwas lowered to 135° C. The mixture was stirred for 2 hours at 135° C. toget the pre-polymer. Nitrogen was vented throughout the aboveprocedures. The pre-polymer was post-polymerized at 120° C. under vacuumfor predetermined time from one day to 3 days to achieve the Poly(1,8-octanediol-citric acid-co-polyethylene oxide). Nitrogen wasintroduced into the reaction system before the polymer was taken outfrom the reaction system. The molar ratios can be altered to achieve aseries of polymers with different properties.

Synthesis of Poly(1,12-dodecanediol-citric acid-co-polyethylene oxide).In a typical experiment, 38.424 g citric acid, 20.234 g1,12-dodecanediol and 40 g polyethylene oxide with molecular weight 400(PEO400)(100 g PEO1000 and 200 g PEO2000 respectively) (molar ratio:citric acid/1,8-octanediol/PE0400=1/0.5/0.5) were added to a 250 ml or500 ml three-neck round-bottom flask, fitted with an inlet adapter andan outlet adapter. The mixture was melted within 15 min by stirring at160-165° C. in silicon oil bath, and then the temperature of the systemwas lowered to 120° C. The mixture was stirred for half an hour at 120°C. to get the pre-polymer. Nitrogen was vented throughout the aboveprocedures. The pre-polymer was post-polymerized at 120° C. under vacuumfor predetermined time from one day to 3 days to achieve thePoly(1,12-dodecanediol-citric acid-co-polyethylene oxide). Nitrogen wasintroduced into the reaction system before the polymer was taken outfrom the reaction system. The molar ratios can be altered to achieve aseries of polymers with different properties.

Synthesis of Poly(1,8-octanediol-citric acid-co-N-methyldiethanoamine)POCM. In a typical experiment, 38.424 g citric acid, 26.321 g1,8-octanediol and 2.3832 g N-methyldiethanoamine (MDEA) (molar ratio:citric acid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or500 ml three-neck round-bottom flask, fitted with an inlet adapter andan outlet adapter. The mixture was melted within 15 min by stirring at160-165° C. in silicon oil bath, and then the temperature of the systemwas lowered to 13520° C. The mixture was stirred for half an hour at120° C. to get the pre-polymer. Nitrogen was vented throughout the aboveprocedures. The pre-polymer was post-polymerized at 80° C. for 6 hours,120° C. for 4 hours without vacuum and then 120° C. for 14 hours undervacuum to achieve the Poly(1,8-octanediol-citricacid-co-N-methyldiethanoamine). Nitrogen was introduced into thereaction system before the polymer was taken out from the reactionsystem. The molar ratios can be altered to citricacid/1,8-octanediol/MDEA=1/0.9510.05.

Synthesis of Poly(1,12-dodecanediol-citricacid-co-N-methyldiethanoamine) PDDCM. In a typical experiment, 38.424 gcitric acid, 36.421 g 1,12-dodecanediol and 2.3832 gN-methyldiethanoamine (MDEA) (molar ratio: citricacid/1,8-octanediol/MDEA=1/0.90/0.10) were added to a 250 ml or 500 mlthree-neck round-bottom flask, fitted with an inlet adapter and anoutlet adapter. The mixture was melted within 15 min by stirring at160-165° C. in a silicon oil bath, and then the temperature of thesystem was lowered to 120° C. The mixture was stirred for half an hourat 120° C. to get the pre-polymer. Nitrogen was vented throughout theabove procedures. The pre-polymer was post-polymerized at 80° C. for 6hours, 120° C. for 4 hours without vacuum and then 120° C. for 14 hoursunder vacuum to achieve the Poly(1,12-dodecanediol-citricacid-co-N-methyldiethanoamine). Nitrogen was introduced into thereaction system before the polymer was taken out from the reactionsystem. The molar ratios can be altered to citricacid/1,12-dodecanediol/MDEA=1/0.95/0.05.

EXAMPLE 2 Materials and Methods used in Preparing and CharacterizingBiodegradable Elastomeric Composites made from POC and Bioceramics

Example 1 describes the production of PDC as well as a number of otherpoly(diol)citrate polymers. In the present Example, there are providedteachings of how to further strengthen and stiffen is thepoly(diol)citrate polymers by incorporating ceramics into theelastomeric polymer matrix.

Materials and Methods

Materials: Hydroxyapatite [Mw: 502.32, Assay>90 (as Ca3 (PO4)₂), 0.5%>75um, 1.4% between 45-75 um, 98.1%<45 um] was purchased from Fluka (St.Louis, Mo., USA). 1,8-octanediol (98%) and citric acid (99.5%) werepurchased from Sigma-Aldrich (St. Louis, Mo., USA). These materials wereused as received. PTFE tubes were purchased from McMaster-CARR, Chicago,USA.

Sample preparation: POC pre-polymer was synthesized according topublished methods [1,5]. Briefly, 0.05 mol of 1,8-octandiol and 0.05 molof citric acid were added to a 100 ml round bottom flask and exposed toa constant flow of nitrogen gas. The mixture was melted under vigorousstirring at 160-165° C. Following melting, the mixture was polymerizedat 140° C. for 1 hr to create a POC pre-polymer. The POC pre-polymer wasmixed with various amounts of HA particles to obtain composites of 40,50, 60, and 65 wt. % HA by mass. Briefly, POC pre-polymer was mixed withthe desired amount of HA powder and placed in PTFE dishes that werepre-warmed to 80° C. The POC-HA mixture was stirred until it becameclay-like, a process that generally took 5-10 hrs depending on the HAcontent. The POC-HA mass was then inserted into PTFE tubes to make rodsor into other PTFE molds designed to meet the dimensional requirementsfor sample mechanical testing protocols or in situ formation of bonescrews. The POC-HA in the mold was then post-polymerized at 80° C. for 3days followed by 120° C. under 2 Pa vacuum for 1 day.

Characterization of the mechanical properties of POC-HA composites: Thefollowing mechanical properties were measured using a Sintech mechanicaltester model 20/G (Triangle Park, N.C. owned by MTS now): 1) bendingstrength (Sb) and modulus (Eb) according to Japanese industrial standard(JIS) K7203, 2) compression strength (Sc) and modulus (Ec) according toJIS K7208, 3) tensile strength (St) and modulus (Et) according to JISK7113, 4) shear strength (Ss)[1,7] and 5) torsional strength (Ts) [18].

All rods used for the mechanical tests were polished with sandpaperbefore measurement. For all mechanical tests, at least 6 samples weretested and the mean values and standard deviations (SD) were calculated.The density of POC and POC-HA composites was measured using theArchimedes principle as previously described [19].

Characterization of morphology of POC-HA composites: SEM was used forobservation of morphology. All POC-HA composites for morphology were thecross section of rods obtained by compression method.

Characterization of the in vitro degradation of POC-HA composites: Thedegradation of POC-HA composite samples (10 mm diameter×2 mm thick) withHA percentages of 40, 50, 60, and 65 wt. % was assessed in vitro in PBS,pH 7.4, at 37° C. for up to 30 weeks under static conditions. Within thePOC-HA composite, only the POC is expected to degrade when incubated inaqueous solution. For comparison purposes, the degradation of POCsamples synthesized under the same conditions as the composites was alsoassessed. PBS was changed as necessary to ensure that the pH did notdrop below 7. Prior to weighing, samples were rinsed with deionizedwater and dried. Mass loss was calculated by comparing the initial mass(Wo) with the mass measured at a given time point (Wt), as shown inEquation 1. The results are presented as means±standard deviation (n=4).$\begin{matrix}{{{Mass}\quad{loss}\quad(\%)} = {\frac{{Wo} - {Wt}}{Wo} \times 100}} & (1)\end{matrix}$

Mineralization of POC-HA composites: Surface mineralization of POC-HAcomposites was assessed in vitro using modified simulated body fluid(SBF)[20]. The SBF consisted of (mmol): Na⁺ (142.0), K+ (4.0),Mg²⁺(1.5), Ca²⁺(5.0), Cl⁻(147), HCO₃ ⁻(4.2), HPO₄ ²⁻(2.0) and SO₄²⁻(0.5) with the pH adjusted to 7.2 usingtris(hydroxymethyl)aminomethane [21]. Briefly, discs (10 mm diameter×2mm thick) of POC-HA composites with HA fraction of 40, 50, 60 and 65 wt.% were immersed in 10 ml of the SBF at 37° C. for up to 15 days. FreshSBF was added every other day to maintain the ionic concentration and pHduring mineralization. The morphology of deposited calcium phosphatecrystals was observed via scanning electron microscopy (SEM) (Hitachi3500 N, EPIC, Northwestern University). The stoichiometric Ca/P molarratio was analyzed by energy dispersive X-ray (EDX).

Evaluation of the cell compatibility of POC-HA composites: POC-HA discs(7.0 mm in diameter×2 mm thick, with 40, 50, 60, and 65 wt. % HA) weresterilized by incubation in 70% ethanol for 30 minutes, washing withsterile PBS (pH 7.4), and UV exposure for 30 minutes. Aftersterilization, samples were washed several times with cell culture mediaprior to placement in the wells of a 48-well tissue culture plate. A 40μl volume of a suspension of human osteoblast cells (HOB) (Cambrex,Pittsburgh, Pa.) (3×10⁵ cells mL⁻¹) was added to each well and incubatedin osteoblast growth medium (OBM and OGM SingleQuots from Cambrex) at37° C. in humidified air and 5% CO2 for up to 14 days. The culturemedium was changed every three days. Samples were fixed with 2.5%gluteraldehyde in PBS for 24 h at 4° C. The morphology of the cells onthe composite samples was observed via SEM.

Statistical methods: Data are expressed as means±standard deviation. Thestatistical significance was calculated using two-tail Student's t-testand analysis of variance (ANOVA) and post-hoc analysis using one-wayanalysis of variation (ANOVA): Newman-Keuls Multiple Comparison Test.P<0.05 was considered as significant differences.

EXAMPLE 3 Preparation and Analysis of Bone Screws made from Compositesof POC and Bioceramics

POC-HA composites having HA fraction of 40 wt. %, 50 wt. %, 60 wt. %, 65wt. % HA were investigated for bone screws. Using compression moldingmethod, POC-HA composites with HA from 40-65 wt. % were compressed intomolds of rods and screws. The POC-HA rods obtained were strong enough tobe further machined into screws. The POC-HA screws from compressionmolding and machining methods were shown in FIG. 1.

Characterization of the mechanical properties of composites. Themechanical property measurements (bending, compression, shear, tensionand torsion) are summarized in Table 1. TABLE 1 Effect of HA fraction onmechanical properties of POC-HA composites POC-HA ρ Sb Eb Sc Ec St Et SsTs Composite (Wt %) (g/cm³) (MPa) (MPa) (MPa) (MPa) (MPa) (MPa) (MPa) (N· m) HA 40 40/60 1.609 33.9 275 32.0 189 7.8 21.4 23.3 22.9 (±0.016)(±5.7) (±80) (±13.0) (±21) (±0.5) (±1.8) (±1.6) (±1.6) HA 50 50/50 1.65337.7 323 64.0 264 7.1 30.2 25.1 24.2 (±0.014) (±4.6) (±63) (±9.4) (±14)(±0.3) (±2.2) (±1.7) (±2.0) HA 60 60/40 1.734 34.7 314 52.6 297 6.4 85.425.9 21.4 (±0.061) (±3.0) (±54) (±11.5) (±41) (±2.0) (±8.8) (±1.6)(±1.9) HA 65 65/35 1.885 41.4 502 74.6 449 9.7 334.8  27.7 27.3 (±0.072)(±3.1) (±40) (±9.0) (±27) (±2.3) (±73.5) (±2.4) (±4.9)Note:ρ: density; Sb: bending strength, Eb: bending modulus, Ss: shearstrength, Sc: compression strength, Ec: compression modulus, St: tensilestrength (rectangular specimens), Et: tensile modulus, Ts: torsionalstrength

(1) Bending strength (Sb) and modulus (Eb): With increasing HA fractionfrom 40 wt. % to 65 wt. %, bending strength (Sb) increased reachingabout 41 MPa at 65 wt. % which was considered as the highest one amongthe composites, and composites showed almost the similar Sb at 40, 50,and 60 wt. %. The bending modulus showed the similar increasingtendencies as bending strength with increasing HA fraction from 40-65wt. %. The highest modulus is 502 MPa at 65 wt. %.

(2) Compression strength (Sc) and modulus (Ec): Compression strength(Sc) greatly increased from 32 to 75 MPa in proportion to HA fractionfrom 40 to 65 wt. % except for 50 wt. %. Compression modulus at 65 wt. %HA is the highest and the values for composites with 40-60 wt. % HA aresimilar to each other. It seems likely that it originated from the HAparticle aggregating in the composites.

(3) Shear strength (Ss): The shear strength (Ss) at 65 wt. % issignificantly higher than that at 40 wt. %, and there is no significantdifference at 40-60 wt. %.

(4) Tensile strength (St) and modulus (Et): Compared with POC, tensilestrength improved greatly for composites with fraction from 40-65 wt. %,and the composite at 65 wt. % has higher tensile strength than those at50 wt. % and 60 wt. %. Modulus increased with increasing HAconcentration from 50-65 wt. % and reached 335 MPa at 65 wt. %, thehighest one among composites.

(5) Torsional strength (Tt): Torsional strength (Tt) did not changesignificantly with increasing HA fraction from 50-60 wt. %. Torsionalstrength at 65 wt. % HA is the highest.

In summary, the POC-HA composites with 65 wt. % HA had the highestbending strength and modulus, compression strength and modulus, tensilestrength and modulus, shear strength and torsional strength.Interestingly, for HA percent of 40-60 wt. % there was no statisticallysignificant difference for most of the mechanical properties. Themechanical properties of all POC-HA composites tested were significantlyincreased relative to POC samples. The density of samples increased withincreasing percentage of HA.

EXAMPLE 4 The Morphology and Other Properties of POC-HA composites

FIG. 2 showed that for the POC composite with 40 wt % HA, HA as nodulesdispersed in POC-HA. However, there were continuous flakes of HAcovering on the surface for the POC composite with 65 wt. % HA,indicating higher amount of HA in this composite.

In vitro degradation of POC-HA composites The mass loss over timeprofiles for POC-HA composites incubated in PBS at 37° C. are shown inFIG. 3. Mass loss is due to the aqueous hydrolytic degradation of POCwithin the composite. The degradation of the POC-HA composite scaffoldswith HA percentages of 40, 50, and 60 was very similar at all timepoints with a total mass loss of approximately 12 wt. % at 20 weeks.POC-HA composite with 65 wt. % HA had the slowest degradation rate witha mass loss of approximately 8.4 wt. % at 20 weeks. For comparisonpurposes, POC samples lost 46 wt. % of their mass by 12 weeks.

Mineralization of POC and PO C-HA composites in SBF SEM showed thatmineralization was not observed on the surface of POC incubated in SBFfor 3 days through 15 days as shown in FIGS. 4 a and b. However, mineralnodules began to form and aggregated on the surface of compositesthroughout the 15 days of incubation in SBF at 37° C. (FIG. 4 c-f).Mineral nodules merged into a continuous covering on most of thesample's surface at 15 days. The composition of the mineral, in terms ofthe molar ratio of Ca/P, was confirmed by the EDX to be 1.5-1.7.Deposition of a large number of calcium phosphates was remarkable on thesurface of composites due to exposure of HA. The osteoconductivity ofPOC-HA composites would be largely predicted.

Evaluation of the cell compatibility of POC-HA composites In order toinvestigate whether HOB can attach and proliferate on the POC-HAcomposites, the morphology of cells seeded on the surface of POC andPOC-HA composites with 40 wt. % and 65 wt. % HA respectively wasobserved by light microscopy (LM) and SEM as shown in FIG. 5. LM showedthat HOB cells can attach, spread out and proliferate on the surface ofPOC cultured from 3 days through 14 days (FIG. 5 a-b). SEM in FIGS. 5 cand e showed that for both POC-HA composites cultured for 3 days, thecells attached and spread out well on the surface of both composites,and layers of the cells on the surface of 65 wt. % POC-HA composite wereobserved. It revealed a good covering on both composites, and bettercell proliferation on 65 wt. % HA composites. While culturing up to 14days, both surfaces were almost completely covered by layers of cells(FIG. 5 d and f). Moreover, the broken cell layers were observed on thesurfaces of both materials in some area due to the thickness of the celllayers. However, HOB cells still remain round aggregating on the POCsurface even culturing for 14 days.

EXAMPLE 5 Further Measurements of Mechanical Properties

(1) Bending strength (Sb) and modulus (Eb). (Japanese IndustrialStandard (JIS) K7203 measurements were determined using a three pointbending test using rods with a range of diameter from 5.0 to 6.5 mm anda length of 30 mm).Sb=8F max L/πD ₃Eb=4L ₃/3 πD ₄(E/Y)where Fmax=maximum force (N); L=support span (mm); D=diameter (mm); andE/Y=gradient of linear portion of stress-strain curve (N/mm).

(2) Compression strength (Sc) and modulus (Ec) (Japanese IndustrialStandard (JIS) K7208 was measured using rods with a range of diameterfrom 5.0 to 6.5 mm and a length of 15-30 mm.Sc=F max/AEc=E/Ywhere Fmax=maximum force (N); A=compressed area (mm2); E/Y=gradient oflinear portion of stress-strain curve (N/mm2).

(3) Shear strength (Ss) (Measured by Suuronen's method [20] at a testingspeed of 10 mm min⁻¹) using rods with a range of diameter from 5.0 to6.5 mm and a length of 20 mm.Ss=F max/2Awhere Fmax=maximum force (N); A=loading area (mm2).

(4) Tensile strength (St) and modulus (Et). (Japanese IndustrialStandard (JIS) K7113 using dog bone shaped samples (26 mm×4 mm×1.6 mm)at testing speed of 10 mm/minSt=F max/AEt=E/YWhere Fmax=maximum force (N); A=transversal cross-sectional area (mm2);E/Y=gradient of linear portion of stress-strain curve (N/mm).

(5) Torsional strength (Ts) The test rod was installed to Sintech 20/GMaterials Testing Machine, using the dumbbell shaped samples having anaverage diameter of 4.7 mm and length of 16.5 mm. The rotating wheel wasturned at a rate of 0.4 rev/min by means of a chain attached to the loadcell. The load pulling the chain was recorded and used for thecalculation of the torque strength according to the equationTs=16F max·R/(πD3)

where Ts=torsion strength (MPa), Fmax=load at fracture (N), R=radius ofthe rotating wheel=38 mm and D=diameter of the rod (mm). TABLE 2 Resultsof ANOVA on Mechanical Properties P-values Sb Eb Sc Ec St Et Ss TsNumber (MPa) (MPa) (MPa) (MPa) (MPa) (MPa) (MPa) (N · m) HA 40 vs N.SN.S P < 0.001 N.S N.S N.S N.S N.S HA 50 HA 40 vs N.S N.S P < 0.05 N.SN.S P < 0.001 N.S N.S HA 60 HA 40 vs P < 0.05 P < 0.001 P < 0.001 P <0.01 N.S P < 0.001 P < 0.01 P < 0.05 HA 65 HA 50 vs N.S N.S P < 0.05 N.SN.S P < 0.001 N.S N.S HA 60 HA 50 vs N.S P < 0.001 N.S N.S P < 0.05 P <0.001 N.S N.S HA 65 HA 60 vs P < 0.05 P < 0.001 P < 0.01 P < 0.05 P <0.01 P < 0.001 N.S P < 0.05 HA 65Note:Sb: bending strength, Eb: bending modulus, Ss: shear strength, Sc:compression strength, Ec: compression modulus, St: tensile strength(rectangular specimens), Et: tensile modulus, Ts: torsional strength.Tensile strength POC vs HA (40-65%): P < 0.001; Tensile modulus POC vsHA (50-65%): P < 0.001: and POC vs HA (40-50%): N.S.

TABLE 3 Results of ANOVA on degradation between POC-HA (40-65 wt. %)composites respectively in 2 and 20 weeks. P-values Time HA40 vs HA40 vsHA50 vs HA40 vs HA50 vs HA60 vs (week) HA50 HA60 HA60 HA65 HA65 HA65 2N.S N.S N.S P < 0.01 P < 0.001 P < 0.001 20 N.S N.S N.S P < 0.05 P <0.05 P < 0.01

TABLE 4 Results of ANOVA on degradation of POC-HA (40-65 wt. %)composites from 2 to 20 weeks. Time P-values (week) HA40 HA50 HA60 HA652 vs 6 P < 0.001 P < 0.001 N.S N.S 2 vs 12 P < 0.001 P < 0.001 N.S N.S 2vs 20 P < 0.001 P < 0.001 N.S P < 0.01 6 vs 12 P < 0.05 P < 0.01 N.S N.S6 vs 20 P < 0.001 P < 0.001 N.S P < 0.01 12 vs 20 P < 0.01 P < 0.05 P <0.05 P < 0.05

REFERENCES

The following references are referred to herein throughout using anumeric identifier. Each of these references is incorporated herein byreference in its entirety.

-   1. PU, R., Absorbable materials in orthopaedic surgery. Ann    Med 1991. 23: p. 109-115.-   2. Yamamuro, T., et al., Bioabsorbable Osteosynthetic Implants of    Ultra-High Strength Poly-LLactide-a Clinical-Study. International    Orthopaedics, 1994. 18(6): p. 332-340.-   3. Habal MB, e., The journal of craniofacial surgery.    1997(Pennsylvania USA: Lippincott-Raven Publishers).-   4. Eppley, B. L. and A. M. Sadove, A comparison of resorbable and    metallic fixation in healing of calvarial bone grafts. Plast    Reconstr Surg, 1995. 96(2): p. 316-22.-   5. Habal, M. B. and W. S. Pietrzak, Key points in the fixation of    the craniofacial skeleton with absorbable biomaterial. Journal of    Craniofacial Surgery, 1999. 10(6): p. 491-499.-   6. Suuronen, R., Biodegradable Fracture-Fixation Devices in    Maxillofacial Surgery. International Journal of Oral and    Maxillofacial Surgery, 1993. 22(1): p. 50-57.-   7. Eppley, B. L., A bioabsorbable poly-L-lactide miniplate and screw    system for osteosynthesis in oral and maxillofacial    surgery—Discussion. Journal of Oral and Maxillofacial Surgery, 1997.    55(9): p. 945-946.-   8. Matsusue, Y., et al., A long-term clinical study on drawn    poly-L-lacticle implants in orthopaedic surgery. Journal of    Long-Term Effects of Medical Implants, 1997. 7(2): p. 119-137.-   9. Bergsma, J. E., et al., Late degradation tissue response to    poly(L-lactide) bone plates and screws. Biomaterials, 1995.    16(1): p. 25-31.-   10. Kikuchi M, C. S.-B., Suetsugu Y, Tanaka J, In vitro tests and in    vivo test developed TCP/CPLA composites. Bioceramics, 1997. 10: p.    407-410.-   11. Reis R l, C. A., Fernandes M H, Correia R N, Bioinert and    biodegradable polymeric matrix composites filled with bioactive    SiO2-3CaOP2O5-MgO glasses and glass-ceramics. Bioceramics, 1997.    10: p. 415-418.-   12. Verheyen, C. C., et al., Evaluation of    hydroxylapatite/poly(L-lactide) composites: mechanical behavior. J    Biomed Mater Res, 1992. 26(10): p. 1277-96.-   13. Verheyen CCPM, K. C., Blieckhogervorst de J M A, Wolke J G C,    Blitterswijk van C A, Groot de K, Evaluation of hydroxyapatite/poly    (L-lactide) composites: physico-chemical properties. J Mater Sci    Mater Med 1993. 4: p. 58-65.-   14. Verheyen, C. C., et al., Hydroxylapatite/poly(L-lactide)    composites: an animal study on pushout strengths and interface    histology. J Biomed Mater Res, 1993. 27(4): p. 433-44.-   15. Jian Yang, A. R. W., Guillermo A. Ameer, Novel citric acid-based    biodegradable elastomers for tissue engineering. Adv. Mater. 2004.    16(6): p. 511-516.-   16. Yang, J., et al., Synthesis and evaluation of poly(diol citrate)    biodegradable elastomers. Biomaterials, 2006. 27(9): p. 1889-98.-   17. Suuronen R, P. T., Taurio R, Törmälä P, Wessman L, Rönkkö K,    Vainionpää S., Strength retention of self-reinforced poly-L-lactide    screws and plates: an in vivo and in vitro study. J Mater Sci Mater    Med 1992. 3: p. 426-431.-   18. T Pohjonen, P.H., P Törmälä, K Koskikare, H Pätiälä, P Rokkanen,    Strength retention of sefreinforced poly-L-lactide screws. A    comparison of compression moulded and machine cut screws. J Mater    Sci Mater Med 1997. 8: p. 311-320.-   19. Yang, J., et al., Fabrication and surface modification of    macroporous poly(L-lactic acid) and poly(L-lactic-co-glycolic acid)    (70/30) cell scaffolds for human skin fibroblast cell culture. J    Biomed Mater Res, 2002. 62(3): p. 438-46.-   20. Zhang, R. and P. X. Ma, Porous poly(L-lactic acid)/apatite    composites created by biomimetic process. J Biomed Mater Res, 1999.    45(4): p. 285-93.-   21. Murphy, W. L. and D. J. Mooney, Bioinspired growth of    crystalline carbonate apatite on biodegradable polymer substrata. J    Am Chem Soc, 2002. 124(9): p. 1910-7.-   22. Shikinami, Y. and M. Okuno, Bioresorbable devices made of forged    composites of hydroxyapatite (HA) particles and poly-L-lactide    (PLLA): Part I. Basic characteristics. Biomaterials, 1999. 20(9): p.    859-877.-   23. Daniels, A. U., M. K. Chang, and K. P. Andriano, Mechanical    properties of biodegradable polymers and composites proposed for    internal fxation of bone. J Appl Biomater, 1990. 1(1): p. 57-78.-   24. Behiri J C, B. M., Khorasani S, Wiwattanadate D, Bonfield W,    Advanced bone cement for longterm orthopaedic implantations In:    Bonfield W, Hastings G W, Tanner K E, editors. Bioceramics, 1991    4(Oxford: Butterworth-Heinemann, UK): p. p. 301-307.-   25. Liu Q, D. W. J. R., Bakker D, Van Blitterswijk C A, Surface    modification of hydroxyapatite to introduce interfacial bonding with    polyactiveTM 70/30 in a biodegradable composites. J Mater Sci Mater    Med., 1996. 7: p. 551-557.-   26. Liu, Q., J. R. de Wijn, and C.A. van Blitterswijk, Covalent    bonding of PMMA, PBMA, and poly(HEMA) to hydroxyapatite particles. J    Biomed Mater Res, 1998. 40(2): p. 257-63.-   27. L L, H., Bioactive ceramics. Ann. New York Acad Sci 1988.    523: p. 54-71.-   28. Li, P. J., et al., Induction and Morphology of Hydroxyapatite,    Precipitated from Metastable Simulated Body-Fluids on Sol-Gel    Prepared Silica. Biomaterials, 1993. 14(13): p. 963-968.-   29. Shikinami, Y. and M. Okuno, Bioresorbable devices made of forged    composites of hydroxyapatite (HA) particles and poly L-lactide    (PLLA). Part II: practical properties of miniscrews and miniplates.    Biomaterials, 2001. 22(23): p. 3197-3211.-   30. Yong Kang, J. Y., Sadiya Khan, Lucas Anissian. Guillermo A.    Ameer A novel biodegradable elastomers for cartilage tissue    engineering. J Biomed Mater Res, 2006. accepted.

1. A composition comprising a composite of: a) a citric acid polyesterhaving the generic formula (A-B-C)_(n), wherein A is a linear aliphaticdihydroxy monomer; B is citric acid, C is a linear aliphatic dihydroxymonomer, and n is an integer greater than 1; and b) a bioceramic usedfor implantable tissue devices, wherein less than 75 wt. % weight ratioof said composition comprises said bioceramic.
 2. The composition ofclaim 1, wherein A is a linear diol comprising between about 2 and about20 carbons.
 3. The composition of claim 1, wherein C is a linear diolcomprising between about 2 and about 20 carbons.
 4. The composition ofclaim 1, wherein both A and C are the same linear diol.
 5. Thecomposition of claim 4, wherein said linear diol is 1, 8, octanediol. 6.The composition of claim 1, wherein A and C are different linear diols.7. The composition of claim 5, wherein said linear aliphatic dihydroxypoly 1,8-octanediol co-citric acid.
 8. The composition of claim 5,wherein said linear aliphatic dihydroxy poly 1,10-decanediol co-citricacid.
 9. The method of claim 1, wherein said bioceramic is selected fromthe group consisting of calcium phosphate bioceramics, alumina-basedbioceramics; zirconia-based bioceramics; silica-based bioceramics, andpyrolytic carbon-based bioceramics.
 10. The method of claim 9, whereinsaid bioceramic is a calcium phosphate bioceramic at a weight percentageof from 30 wt. % to about 75 wt. % of the total weight of thecomposition.
 11. The method of claim 10, wherein said bioceramic ishydroxyapatite (HA) at a weight percentage of between about 40 wt. %±5wt. % to about 70 wt. %±5 wt. % HA to 35 wt. %±5 wt. % to about 25 wt.%±5 wt. % citric acid polyester.
 12. The composition of claim 1, whereinsaid composition comprises at least 40 wt. % HA but less than 10 wt. %HA.
 13. The composition of claim 12, wherein said composition comprisesat least 45 wt. % HA.
 14. The composition of claim 12, wherein saidcomposition comprises at least 50 wt. % HA.
 15. The composition of claim12, wherein said composition comprises at least 60 wt. % HA.
 16. Thecomposition of claim 12, wherein said composition comprises at least 65wt. % HA.
 17. The composition of claim 12, wherein said compositioncomprises at least 70 wt. % HA.
 18. The composite of any of claims 1 to17 wherein said composite has a bending strength of from about 33.9 toabout 41.4 MPa.
 19. The composite of any of claims 1 to 17 wherein saidcomposite has a compression strength of from about 32 to about 75 MPa.20. The composite of any of claims 1 to 17 wherein said composite has atensile strength of from about 6 to about 10 MPa.
 21. The composite ofany of claims 1 to 17 wherein said composite has a shear strength offrom about 23 to about 28 MPa.
 22. The composite of any of claims 1 to17 wherein said composite has a bending modulus of from about 0.275 toabout 0.502 GPa.
 23. The composite of any of claims 1 to 17 wherein saidcomposite has a compression modulus of from about 0.19 to about 0.45GPa.
 24. The composite of any of claims 1 to 17 wherein said compositehas a compression modulus of from about 0.02 to about 0.34 GPa.
 25. Thecomposition of claim 1, wherein said composition further comprises apolymer is selected from the group consisting of poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(trimethylene carbonate), and polyester amide.26. The composition of claim 1, wherein said composition is molded intoan orthopedic fixation device.
 27. The composition of claim 26, whereinsaid orthopedic fixation device is selected from the group consisting ofa bone screw, a bone pin, a bone rod, and a bone plate.
 28. Anartificial bone, wherein said bone is comprised of a composition of anyof claims 1-17.
 29. A substrate for use in an implantable devicecomprising a composition of any of claims 1-17, formulated into anorthopedic fixation device or an artificial bone structure.
 30. Thesubstrate of claim 29, wherein said substrate further comprises asurface modification to facilitate implantation of said device with adecreased risk of implant rejection.
 31. A method of producing animplantable device, comprising: a. preparing a composition according toany of claims 1-17; and b. molding the composition of (a) into anorthopedic fixation device or an artificial bone for implantation. 32.The method of claim 31, wherein said fixation device is selected fromthe group consisting of a bone screw, a bone pin, and a bone plate. 33.An implantable device comprising a polymer composition of any of claims1-17.
 34. A composition comprising a composite of: a) a citric acidpolyester having the generic formula (A-B-C)_(n), wherein A is a linearaliphatic dihydroxy monomer; B is citric acid, C is a linear aliphaticdihyroxy monomer, and n is an integer greater than 1; and b) abioceramic used for implantable tissue devices, wherein at least 30 wt.% weight ratio of said composition comprises said bioceramic.